1. Field of the Invention
The present invention concerns a device for determining the presumed position of an event inducing a signal, in photodetectors, this position being, for example, located in relation to the photodetector assembly.
The invention applies in particular to determination of the position of an event from signals supplied by photomultipliers equipping a gamma camera, the position being located in relation to the photomultipliers themselves. Gamma camera means a camera sensitive to gamma (.gamma.) radiation. Such cameras are used notably for the purposes of medical imaging.
2. Discussion of the Background
At the present time, the majority of gamma cameras used in nuclear medicine are cameras operating according to the principle of Anger type cameras. This subject can be referred to U.S. Pat. No. 3,011,057.
Gamma cameras make it possible in particular to visualize the distribution, in an organ, of molecules labelled with a radioactive isotope previously injected into the patient.
The structure and operation of a known gamma camera are described and summarized below with reference to the accompanying FIGS. 1, 2A and 2B.
FIG. 1 shows a detection head 10 of a gamma camera disposed opposite an organ 12 containing molecules labelled with a radioactive isotope.
The detection head 10 has a collimator 20, a scintillator crystal 22, a light guide 24 and a plurality of photomultiplier tubes 26 juxtaposed so as to cover one face of the light guide 24. The scintillator is, for example, a crystal of NaI (Tl).
The function of the collimator 20 is to select from among all the gamma radiations 30 emitted by the organ 12 those which reach the detection head substantially at normal incidence. The selective nature of the collimator makes it possible to increase the resolution and clarity of the image produced. However, the increase in resolution is made to the detriment of the sensitivity. By way of example, for around 10,000 .gamma. photons emitted by the organ 12, one single photon is actually detected.
The .gamma. photons which have passed through the collimator reach the scintillator crystal 22 where nearly every .gamma. photon is converted into a plurality of light photons. In the remainder of the text, event designates each interaction of a gamma photon with the crystal, causing a scintillation.
The photomultipliers 26 are designed to emit an electrical pulse proportional to the number of light photons received from the scintillator for each event.
So that a scintillation event can be localized more precisely, the photomultipliers 26 are not placed directly side by side with the scintillator crystal 22 but are separated from the latter by the light guide 24.
The photomultipliers emit a signal whose amplitude is proportional to the total quantity of light produced in the scintillator by gamma radiation, that is to say proportional to its energy. However, the individual signal from each photomultiplier also depends on the distance which separates it from the point of interaction 30 of the gamma radiation with the material of the scintillator. This is because each photomultiplier delivers a current pulse proportional to the light flux it has received. In the example of FIG. 1, small graphs A, B, C show that photomultipliers 26a, 26b and 26c situated at different distances from a point of interaction 30 deliver signals with different amplitudes.
The position of the point of interaction 30 of a gamma photon is calculated in the gamma camera from signals coming from the photomultiplier assembly by performing a barycentric weighting of the contributions of each photomultiplier.
The principle of barycentric weighting as implemented in Anger type cameras emerges more clearly on referring to the accompanying FIGS. 2A and 2B.
FIG. 2A shows the electrical wiring of a detection head 10 of a gamma camera, which connects this camera to an image forming unit. The detection head has a plurality of photomultipliers 26.
As shown in FIG. 2B, each photomultiplier 26 of the detection head is associated with four resistors denoted RX.sup.-, RX.sup.+, RY.sup.- and RY.sup.+. The values of these resistors are specific to each photomultiplier and depend on the position of the photomultiplier in the detection head 10.
The resistors RX.sup.-, RX.sup.+, RY.sup.- and RY.sup.+ of each photomultiplier are connected to the output 50 of the said photomultiplier, represented in FIG. 2B by a current generator symbol. They are moreover respectively connected to common collector lines denoted LX.sup.-, LX.sup.+, LY.sup.- and LY.sup.+, in FIG. 2A.
The lines LX.sup.-, LX.sup.+, LY.sup.- and LY.sup.+ are in turn connected respectively to analog integrators 52X.sup.-, 52X.sup.+, 52Y.sup.- and 52Y.sup.+, and, by means of the latter, to analog-to-digital converters 54X.sup.-, 54X.sup.+, 54Y.sup.- and 54Y.sup.+. The output of the converters 54X.sup.-, 54X.sup.+, 54Y.sup.- and 54Y.sup.+ is taken to a digital operator 56. The lines LX.sup.-, LX.sup.+, LY.sup.- and LY.sup.+ are furthermore connected to a common path, referred to as the energy path. This path also has an integrator 57 and an analog-to-digital converter 58 and its output is also taken to the operator 56.
By virtue of the device of FIG. 2, the position of the interaction is calculated according to the following equations (U.S. Pat. No. 4,672,542): ##EQU1##
in which X and Y indicate the coordinates, along two orthogonal directions, of the position of the interaction on the crystal and in which X.sup.+, X.sup.-, Y.sup.+, Y.sup.- indicate respectively the weighted signals delivered by the integrators 52X.sup.+, 52X.sup.-, 52Y.sup.+, 52Y.sup.-.
The values of X and Y, as well as the total energy E of the gamma ray which has interacted with the crystal, are produced by the digital operator 56. These values are next used for constructing an image as described, for example, in the document FR-2 669 439.
The calculation of the position of the interaction is marred by an uncertainty related to the statistical Poisson fluctuations of the number of light photons and the number of photoelectrons produced for each event, that is to say for each gamma photon detected. The higher the number of photons or photoelectrons, the smaller is the standard deviation of the fluctuation. Because of this phenomenon, the light should be collected as carefully as possible. The intrinsic spatial resolution of the camera is characterised by the mid-height width of the distribution of the positions calculated for one and the same collimated point source placed on the scintillator crystal.
For gamma rays with an energy of 140 keV, the resolution is generally of the order of 3 to 4 mm.
The energy of a detected gamma photon is calculated by summing the contributions of all the photomultipliers which have received light. This is also marred by a statistical fluctuation. The resolution energy-wise of the camera is characterised by the ratio of the mid-height width of the distribution of the calculated energies to the mean value of the distribution, for one and the same source.
The resolution energy-wise is generally of the order of 9 to 11% for gamma rays with an energy of 140 keV.
Finally, an Anger type gamma camera has the advantage of making it possible to calculate in real time the barycentre of the signals from photomultipliers with very simple means.
This is because the system described previously has a limited number of components. Moreover, the resistors used to inject the signal from the photomultipliers into the collector lines are very inexpensive.
Such a camera has however also a major drawback, which is a reduced counting rate. Counting rate means the number of events, that is to say interactions between a .gamma. photon and the scintillator, which the camera is capable of processing per unit of time.
One of the limitations of the counting rate is due notably to the fact that the camera is incapable of processing two events taking place substantially simultaneously at distinct points of the scintillator crystal.
This is because simultaneous but geometrically distinct events give rise to electrical signals which stack up in the collector lines LX.sup.-, LX.sup.+, LY.sup.- and LY.sup.+ and which can no longer be distinguished. These events are also "lost" for the formation of an image.
The limitation of the counting rate is not too important a constraint in the traditional medical imaging techniques. This is because, as indicated above, the collimator stops a very large number of gamma rays and only a small number of events are actually detected.
Gamma cameras are however also used in two other medical imaging techniques where the limitation of the counting rate is a crippling constraint.
These techniques are the so-called "attenuation equalisation by transmission" and "coincidence PET (Positron Emission Tomography)" techniques.
The attenuation equalisation by transmission technique consists of taking into account, at the time of forming a medical image, the attenuation belonging to the tissue of the patient surrounding the organ examined. In order to know this attenuation, a measurement is made of the transmission of gamma radiations to a gamma camera through the body of the patient. To that end the patient is made to take a position between a highly active external source and the detection head of the gamma camera. Thus, at the time of measuring the transmitted radiation, a high number of events take place in the scintillator crystal. The high number of events per unit of time also increases the probability of having a number of substantially simultaneous events. A conventional Anger type camera then proves to be inappropriate.
The PET technique consists of injecting into the patient an element such as F.sup.18 capable of emitting positrons. The annihilation of a positron and an electron releases two .gamma. photons emitted in opposite directions and having an energy of 511 keV. This physical phenomenon is taken advantage of in the PET imaging technique. In this technique, use is made of a gamma camera with at least two detection heads disposed on either side of the patient. The detection heads used are not equipped with a collimator. This is because electronic processing of the information, referred to as coincidence processing, makes it possible to select, from among the events, those which coincide timewise, and to thus calculate the trajectory of the gamma photons.
The detection heads are therefore subjected to high gamma radiation fluxes. The conventional Anger type gamma cameras have a counting rate which is generally too limited for such an application.
For information only, an Anger type gamma camera can normally operate with a detection of 1.10.sup.5 events per second, whereas in PET imaging at least 1.10.sup.6 events per second are necessary for normal operation.
Another limitation of Anger type gamma cameras, described above, is due to the fact that the calculation of the barycentre of an event is definitively fixed by the construction of the detection head and notably by the choice of the resistors RX.sup.-, RX.sup.+, RY.sup.-, RY.sup.+ for each photomultiplier. Similarly, the calculation of the energy is fixed by the wiring of the photomultipliers on to a common path (the energy path).